One in eight women may develop breast cancer in her life time, and mammography has largely been credited for the early detection of breast cancer that leads to an early therapy and a five year survival rate of well beyond 90%.
Although, as discussed below, more advanced digital technologies have been developed to improve mammography image quality, there are three limitations associated with mammography that call for improvements. First, mammograms miss up to 20% of breast cancers that are present as false negative. Second, in some cases a mammogram appears abnormal, yet there is no breast cancer, thus result in a false positive. Third, the risk of radiation induced carcinoma for woman in general before age 50 due to the statistics that mammography may not gain sufficient benefit from its use. Also most patients complain on the procedure of breast compression.
In more practical considerations of instrument design, soft tissues have very weak shadow contrast from X-ray attenuation contrast (XAC), and mammography allows only ˜5% of the X-ray beam to reach the detector while 95% of fluence is absorbed by the compressed breast in order to deliver a useful shadow image from a weak XAC. This XAC approach limits the beam energy for mammography at 14±3 keV, which can be delivered by a Coolidge tube using a rotating anode at a tube bias of 24±4 kVp. The anode metal is coated with Rh in order to avoid any line-emission which cannot be part of the variable of bias necessary to adjust for a range of tissue thickness and density. Also because of the interference to the imager from wide-angled Compton scattering, the use of moving Bucky grid with a narrow cone beam angle to cover the breast restricted the minimal length of X-ray beam path to be long and the instrument less efficient. Thermal load of the mammo-tube at 4.5-6.5 kW implies the use of a rotating anode to spread the heat from an e-beam focal point.
Some of the digital advancements alluded to above are described in US Patent Application Publication 2015/0139390 to Bellazzini (the contents of which are incorporated herein by reference). For example, Bellazzini describes the use of digital X-ray sensors comprising a conversion layer in the form of an amorphous coating, normally made of Amorphous Selenium or of Cesium iodide, and an integration panel, i.e. a collection layer, that has a TFT pixel structure (Thin Film Transistor). The conversion layer serves for transforming into an electric charge the photons of an X-ray beam that has travelled across an irradiated sample. This may occur directly or indirectly, as in the case of amorphous Selenium and of Cesium iodide; respectively. The total charge obtained by the conversion during an X-ray exposure builds up in the pixels of the integration panel. Once the exposure has been completed, the amount of charge accumulated in each pixel is read. More in detail, an image acquisition electronics is provided that comprises an analog-to-digital converter arranged at the boundary of the integration panel (or at a second chip to be stacked beneath the sensor chip). The analog-to-digital converter changes the overall charge accumulated in each pixel into an electric voltage, i.e. into a number that is proportional to the overall radiation that has travelled across the sample at each pixel of the integration panel. These numbers can be converted into a radiographic image in which the contrast depends upon the overall radiation that is accumulated in each pixel.
Bellazzini also describes a so-called “photon-counting” technique that is also described, for example, in U.S. Pat. No. 8,680,474 to Soh, et al (the contents of which are incorporated herein by reference). Using a photon-counting technique, photons can be counted one-by-one, and ranked into a plurality of channels, thus obtaining a “film grade” resolution, i.e. a resolution that is comparable with the resolution allowed by high-resolution radiographic plates. In particular, hybrid detectors exist that are known as Medipix and that are provided with an ASIC for carrying out a photon-counting procedure. These hybrid detectors comprise discriminators associated with event counters that are used in such a way that the image acquisition electronics counts only events, i.e. acquisitions of photons that fall in a predetermined energy window. This way, an X-ray imaging technique is obtained that has spectroscopic features. A more recent device, known as Medipix-3, has a finer energy resolution thanks to a real-time charge share correction. Medipix-3 also comprises multiple pixel counters that can be used in different operation modes. This allows a continuous detection, and up to eight energy thresholds can be obtained.
In the Medipix device, like in other devices, the collection layer is implemented by CMOS technology, which is a low-power consumption device, i.e. about a few Watts, that can be fabricated at a low-cost. As has been developed for optical imagers, the CMOS imagers need to be compared with film resolution at 60 line-pairs per millimeter (lp/mm). A large number of pixels per chip, e.g. about 107, is required for a CMOS imager to compare with x-ray films. For a radiological digital panel with much larger pixel sizes the resolution is only approximately 6 lp/mm.
Notwithstanding these digital manipulations, there is still an urgent need for an apparatus and method that enables enhanced intrinsic resolution mammography and provides improvements over the x-ray attenuation contrast (XAC) imaging discussed above.